Opposite to conventional radiography wherein an intensifying luminescent phosphor screen directly emits luminescent radiation and wherein said screen is not a storage medium, radiation image recording and reproducing techniques utilizing a radiation image storage panel, referred to as the stimulable phosphor screen, sheet or panel, are provided with a stimulable phosphor. With radiation image recording and reproducing techniques, the stimulable phosphor of the radiation image storage panel is caused to absorb radiation, which carries image information of an object or which has been radiated out from a sample. Said stimulable phosphor is exposed to stimulating rays, such as visible light or infrared rays, which causes the stimulable phosphor to emit light in proportion to the amount of energy stored thereon during its irradiation exposure.
The emitted fluorescent light is then photoelectrically detected in order to obtain an electric signal. The electric signal is further processed, and the processed electric signal is utilized for reproducing a visible image on a recording material. This way of working, making use of storage phosphor sheets or panels as an intermediate storage medium is also called “computed radiography”.
As in radiography it is important to have excellent image quality for the radiologist to make an accurate evaluation of a patient's condition, important image quality aspects are image resolution and image signal-to-noise ratio.
For computed radiography signal-to-noise ratio depends on a number of factors.
First, the number of X-ray quanta absorbed by the storage phosphor screen is important. Signal-to-noise ratio will be proportional to the square-root of the number of absorbed quanta.
Second, the so-called fluorescence noise is important. This noise is caused by the fact that the number of photostimulated light (PSL) quanta detected for an absorbed X-ray quantum is small. Since a lot of the PSL light is lost in the detection process in computer radiography, fluorescence noise has an important contribution to the signal-to-noise ratio. It is important that, on the average, at least 1 photon is detected for every absorbed X-ray quantum. If this is not the case, many absorbed X-ray quanta will not contribute to the image and signal-to-noise ratio will be very poor.
This situation is most critical in mammography, where X-ray quanta are used with low energy. Softer X-rays will give rise to less PSL centres and, therefore, to less PSL photons than harder X-rays.
In computed radiography, a number of PSL centres are created by the absorbed X-ray quanta. Not all PSL centres are stimulated in the read-out process, because of the limited time available for pixel stimulation and because of the limited laser power available. In practice, only about 30% of the PSL centres is stimulated to give rise to a PSL photon. Since these photons are emitted and scattered in all directions, only 50% of the PSL photons are emitted at the top side of the storage phosphor screen, where they can be detected by the detection system. The emitted PSL photons are guided towards the detector by a light guide. This light guide may consist of an array of optical fibres, that forms rectangular detection area above the storage phosphor screen and has a circular cross-section at the detector side. This type of light guide has a numerical aperture of only 30%, which means that only 1 out of 3 of the emitted PSL photons is guided to the detector. In between the light guide and the detector a filter is placed, which stops the stimulation light reflected by the storage phosphor screen and transmits the PSL light emitted by the screen. This filter also has a small absorption and reflection of PSL light and transmits only ca. 75% of the PSL photons. In computed radiography a photomultiplier is commonly used to transform the PSL signal into an electric signal. At 440 nm the photomultiplier has a quantum efficiency of ca. 20%. This means that only 1 out of 5 PSL quanta that reach the photomultiplier are detected.
In summary, for 1,000 PSL centres that are created in the storage phosphor screen, only 1,000×0.3×0.5×0.3×0.75×0.2 or 6.75 PSL photons are detected. If it is required that every X-ray quantum gives rise to at least 1 detected PSL photon, therefore, the number of PSL centres created by an X-ray quantum should be sufficiently large. Or, conversely, the X-ray energy required to create a PSL-centre should be sufficiently small.
In mammography, a common setting of the X-ray source is at 28 kVp. This leads to an X-ray spectrum, where the average energy of an X-ray quantum is of the order of 15 keV. For an X-ray quantum with this energy, in order to give rise to at least 1 detected PSL photon, the energy needed to create a PSL centre should be less than 15,000×6.75/1,000=100 eV.
Further it is well-known that fine detail visualization, high-resolution high-contrast images are required for many X-ray medical imaging systems and particularly in mammography. The resolution of X-ray film/screen and digital mammography systems is currently limited to 20 line pairs/mm and 10 line pairs/mm, respectively. One of the key reasons for this limitation is associated with the phosphor particle size in the currently used X-ray screens.
In particular, light scattering by the phosphor particles and their grain boundaries results in loss of spatial resolution and contrast in the image. In order to increase the resolution and contrast, scattering of the visible light must be decreased. Scattering can be decreased by reducing the phosphor particle size while maintaining the phosphor luminescence efficiency. Furthermore, the X-ray to light conversion efficiency, the quantum detection efficiency (e.g. the fraction of absorbed X-rays convertable to light emitted after stimulation) and the screen efficiency (e.g. the fraction of emitted light escaping from the screen after irradiation with stimulating rays) should not be affected in a negative way by the reduction of the phosphor particle size. As a particular advantage the computed radiographic recording and reproducing techniques presented hereinbefore show a radiation image containing a large amount of information, obtainable with a markedly lower dose of radiation than in conventional radiography.
For clinical diagnosis and routine screening of asymptomatic female population, screen-film mammography today still represents the state-of-the-art technology for early detection of breast cancer. However, screen-film mammography has limitations which reduce its effectiveness. Because of the extremely low differentiation in radiation absorption densities in the breast tissue, image contrast is inherently low. Film noise and scatter radiation further reduce contrast making detection of microcalcifications difficult in the displayed image. So e.g. film gradient must be balanced against the need for wider latitude.
Digital radiography systems can be broadly categorized as primary digital and secondary digital systems. Primary digital systems imply direct conversion of the radiation incident on a sensor into usable electric signals to form a digital image. Secondary digital-systems, on the other hand, involve an intermediary step in the conversion of radiation to a digital image. For example, in digital fluoroscopy, image intensifiers are used for intermediary conversion of X-rays into a visible image that is then converted to a digital image using a video camera. Similarly, digital X-ray systems using photostimulated luminescence (PSL) plates, first store the virtual image as chemical energy. In a second step, the stored chemical energy is converted into electric signals using a laser to scan the PSL plate to form a digital image.
Furthermore, various schemes using silicon photodiode arrays in scanning mode for digital radiography systems have been employed. However, these photodiode arrays require intermediate phosphor screens to convert X-rays into visible light, because of the steep fall-off in quantum efficiency (sensitivity) of the arrays at energies above 10 keV.
A preferably employed stimulable phosphor, embedded in a phosphor plate, is a phosphor which absorbs not only a radiation having a wavelength lower than 250 nm but also visible or ultraviolet light in the wavelength region of 250 to 400 nm, and further gives a stimulated emission of a wavelength in the range of 300 to 500 nm when it is irradiated with stimulating rays in the wavelength range of 400 to 900 nm.
Examples of well-known, frequently used stimulable phosphors include divalent europium activated phosphors (e.g., BaFBr:Eu, BaFBrI:Eu) or cerium activated alkaline earth metal halide phosphors and cerium activated oxyhalide phosphors, as well as e.g. a phosphor having the formula of YLuSiO5:Ce,Zr.
In the present invention it is envisaged to randomly use screens containing either divalent europium activated alkali halide type phosphor screens, wherein said halide is at least one of chloride, bromide and iodide or a combination thereof or divalent europium activated alkaline earth metal phosphor screens wherein said halide is at least one of fluoride, chloride, bromide and iodide or a combination thereof. Most preferred is random use of divalent europium activated CsX type phosphor screens, wherein said X represents Br or a combination of Br with at least one of Cl and I, as Br(Cl), Br(I) or Br(Cl,I) and bariumfluorohalide phosphor screens wherein the phosphor is of the (Ba,MII)FX′:Eu type, wherein MII is an alkaline earth metal and wherein X′ is Cl, Br and/or I.
Crystalline divalent europium activated alkali halide phosphor screens advantageously have CsBr:Eu2+ storage phosphor particles, in binderless layers in the form of cylinders (and even up to a needle-shaped form) wherein said cylinder has an average cross-section diameter in the range from 1 μm to 30 μm (more preferred: from 2 μm up to 15 μm), an average length, measured along the casing of said cylinder, in the range from 100 μm up to 1000 μm (more preferred: from 100 μm up to 500 μm) as has e.g. been described in EP-A 1 359 204. Such block-shaped, prismatic, cylindrical or needle-shaped phosphors, whether or not obtained after milling, are, in another embodiment, coated in a phosphor binder layer.
According to another embodiment of the present invention said stimulable phosphors are (Ba,MII)FX′:Eu type phosphors, wherein MII is an alkaline earth metal and wherein X′ is Cl, Br and/or I. In a preferred embodiment, said MII is Sr2+. Powder phosphor screens that are advantageously used in the system of the present invention have europium activated alkaline earth metal halide phosphor screens containing Ba(Sr)FBr:Eu2+ storage phosphor particles, dispersed in a binder medium in their corresponding storage phosphor layers.
The recorded image itself is reproduced by stimulating the exposed photostimulable phosphor screens by means of stimulating radiation and by detecting the light that is emitted by the phosphor screen upon stimulation and converting the detected light into an electric signal representation of the radiation image.
In a specific embodiment light emitted by the phosphor screen upon stimulation is detected by means of an array of charge coupled devices. In order to obtain a good collection efficiency the light emitted by the phosphor screen upon stimulation is guided by means of a light guide onto the array of charge coupled devices. In one embodiment this light guide is implemented in the form of a fibre optic plate (FOP). A FOP plate consists of a number of juxtaposed optical fibres that together form a two-dimensional light guiding array. The first dimension of the array corresponds with the length of a scan line on the photostimulable phosphor screen while the second dimension covers the width of the array of transducer elements. In this way the light emitted when stimulating a scan line on the photostimulable phosphor screen is guided onto the array of transducer elements in a point-by-point like fashion. The light the phosphor screen is exposed to in order to be stimulated should be separated from the light emitted by the screen upon stimulation. An easy way to separate stimulating light and emission light is to make use of an optical filter in between the light input face of the fibre optic plate and the phosphor screen. Colored glass filters are widely used for this purpose.
An optimized resolution is obtained when the fibre optic plate is in close contact with the phosphor screen. Colored glass filters however are rather thick, so that provision of a colored glass filter in between the fibre optic plate and the phosphor screen is incompatible with the requirement of having close contact between the fibre optic plate and screen.
Whatever stimulable or storage phosphor screen is applied in medical diagnostic imaging, light emitted by the phosphor screen upon stimulation should be separated from stimulation light. When use is made of a storage phosphor panel containing e.g. a CsBr:Eu phosphor, the stimulating light source is a light source emitting light in the range of 600 to 800 nm and the filter should absorb the laser light to an extent as high as possible, while at the same time absorption by the same filter of the light emitted upon stimulation, having maximum emission of radiation at 440 nm, should be minimized.
An optimized optical density of the filter in the stimulation wavelength range should at least have a density value of 6, while the transmission in the emission wavelength range should at least exceed 50%, resulting in a density equal to or less than 0.30. An optical density of 6 means that the laser light is attenuated with a factor of 106, or otherwise expressed that an absorption of 99.9999% is attained. To achieve these specifications by means of a glass filter such as a BG 39 Schott® filter, the thickness of the filter should at least be 5 mm, being extremely thick and laying burden on sharpness.
In case of a read-out apparatus as described higher wherein the light emitted by the phosphor screen is guided to the array of transducer elements via a FOP, the gap between the input face of the FOP and the phosphor screen can only be approximately 100 μm in order to provide the desired high resolution. It is thus clear that a glass filter is not suitable for this application. When the FOP is replaced by an array of microlenses or a selfoc array, this gap would attain a value in the range of 2 to 3 mm. Even in this case use of a glass filter as described higher would provide insufficient sharpness.
From a point of view of practical use, as already suggested hereinbefore, different stimulable phosphor screens or panels are desired, all of them, in at random order, to give stimulated emission in the wavelength range of 300–500 nm when excited with stimulating rays in the wavelength range of 500–850 nm. Preventing the 500–850 nm light from reaching the detector is particularly important when the detector is a CCD having the highest quantum efficiency in the red region. The stimulation light can only be filtered away when the wavelength of the light emitted upon stimulation is quite different from the green or red stimulation light, i.e., that there is no or a only negligable overlap between the stimulation radiation spectrum and the stimulated emission radiation spectrum. In favor of customer-friendly handling or manutention in a medical radiographic environment, wherein a lot of phosphor plates or panels are exposed and read out (processed) one after another, even if processed in an at random order, it is recommended that detection of the blue light, emitted after photostimulation, proceeds with filters transmitting blue light for all screens or panels, without the need to change filters inbetween consecutive readings. Use of only one, same filter for all of the different plates scanned in one, same scanning unit, in applications requiring an optimized image quality as well as in applications requiring ordinary image quality, would be highly desired, more particularly in favor of cost reduction. Use of only one scanner would be highly appreciated for different types of plates.
In one aspect, as in EP-A 1 065 525, a system for reading a radiation image from a phosphor screen comprises a specific divalent europium activated cesium halide phosphor, provided with filtering means comprising a dye in order to prevent light emitted by said source of stimulation light from being detected by transducer elements and is restricted to one phosphor type. As the dye has an absorption spectrum with an absorption peak falling within the range of 600 to 800 nm, with a maximum of said peak attaining a value corresponding with at least 99% absorption and with an absorption in the range of 400 to 500 nm of less than 25%, this is not sufficient in order to bring a solution for the problems set out hereinbefore.